This invention relates generally to methods and apparatus for providing increased resolution and/or additional imaging slices for computed tomography imaging systems, and especially to methods and apparatus for upgrading resolution and/or imaging slices coverage of existing computed tomographic imaging systems.
In at least one known computed tomography (CT) imaging system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the xe2x80x9cimaging planexe2x80x9d. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile.
In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a xe2x80x9cviewxe2x80x9d. A xe2x80x9cscanxe2x80x9d of the object comprises a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector.
In an axial scan, the projection data is processed to construct an image that corresponds to a two dimensional slice taken through the object. One method for reconstructing an image from a set of projection data is referred to in the art as the filtered back projection technique. This process converts the attenuation measurements from a scan into integers called xe2x80x9cCT numbersxe2x80x9d or xe2x80x9cHounsfield unitsxe2x80x9d, which are used to control the brightness of a corresponding pixel on a cathode ray tube display.
To reduce the total scan time required for multiple slices, a xe2x80x9chelicalxe2x80x9dscan may be performed. To perform a xe2x80x9chelicalxe2x80x9d scan, the patient is moved in the z-axis synchronously with the rotation of the gantry, while the data for the prescribed number of slices is acquired. Such a system generates a single helix from a fan beam helical scan. The helix mapped out by the fan beam yields projection data from which images in each prescribed slice may be reconstructed. In addition to reducing scan time, helical scanning provides other advantages such as better use of injected contrast, improved image reconstruction at arbitrary locations, and better three-dimensional images.
The axis of rotation of the rotating gantry on which the detector and x-ray source rotate is referred to as a z-axis. The detector array is considered as having a z-direction defined as a direction parallel to the z-axis of the rotating gantry. When a patient is scanned, the z-direction of the detector array is usually at least approximately aligned with the patient""s spine.
A single-slice detector array is used in at least one known CT imaging system. Referring to the representation of a portion of a single slice detector array 18 of FIG. 3, only one detector row extends in a direction transverse to the z-direction. At least one other known CT imaging system utilizes a multi-slice detector array having more than one row, for example, four rows or sixteen rows, each row extending in a direction transverse to the z-direction. The CT imaging systems referred to here have detectors arrays in which the rows are arranged in one or more arcs transverse to the z-axis. It is thus convenient to define an x-direction as a direction along the arc of the detector. Detectors are often represented by flat, two-dimensional projections of their target surfaces, with one dimension representing the z-direction and the other dimension representing the x-direction. This convention is used in many of the figures in this description.
In known CT imaging systems utilizing the embodiment of prior art detector array 18 of FIG. 3, the data acquisition system and detector array of the prior art imaging system do not provide a trade-off between resolution and number of slices. In all cases, exactly one slice of data is received from the detector at a time. Some additional flexibility is provided in another known imaging system that utilizes a multislice detector having rows not all having the same thickness in the z-direction. However, the bandwidth of the data acquisition system in this system is able to process no more than a fixed maximum number of slices that is less than the number of rows of detector elements 20 in detector array 18. For example, in imaging systems having sixteen row detector arrays, only four slices of attenuation data can be acquired at one time. When using more than four rows of detector elements 20, the outputs of detector elements 20 in selected adjoining rows of detector array 18 are combined in the z-direction to keep the number of slices, and hence, the data bandwidth of the imaging system at or below its maximum limit.
It would thus be desirable to provide existing clinical computed tomography (CT) imaging with enhanced productivity and new applications by increasing the range of imaging combinations available. True isotropic and/or high-resolution volumetric scanning, for example, would enable new cardiac, interventional and screening applications as well as improved image quality for at least heads and inner ears with an increased productivity.
There is thus provided, in one embodiment of the present invention, a method for changing a number of image slices and/or in-plane resolutions available in an existing imaging system having a radiation source, an existing detector array having an x-direction and a z-direction and configured to acquire attenuation measurements of an object between the radiation source and the existing detector array, an image reconstructor configured to reconstruct an image of the object from attenuation data, and a communication path between the existing detector array and the image reconstructor; the communication path including a data acquisition system coupled between the existing detector array and the imaging reconstructor, the communication path also having a maximum bandwidth limit. The method includes steps of: replacing the existing detector array with a replacement detector array having either narrower detector cells in the x-direction than the existing detector array, a greater number of detector cells than that of the existing detector array, or both; and selecting an in-plane resolution of the replacement detector array in accordance with the maximum bandwidth limit of the communication path.
Isotropic and/or high resolution volumetric scanning is achieved in embodiments of the present invention by using replacement detectors capable of imaging a larger number of slices and/or having higher resolution in the x- and/or z-direction.